The use of positron emission tomography is growing in the field of medical imaging. In PET imaging, a radiopharmaceutical agent is introduced into the object to be imaged via injection, inhalation, or ingestion. After administration of the radiopharmaceutical, the physical and bio-molecular properties of the agent will concentrate at specific locations in the human body. The actual spatial distribution of the agent, the intensity of the region of accumulation of the agent, and the kinetics of the process from administration to eventually elimination are all factors that may have clinical significance. During this process, a positron emitter attached to the radiopharmaceutical agent will emit positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc.
The radionuclide emits positrons, and when an emitted positron collides with an electron, an annihilation event occurs, wherein the positron and electron are destroyed. Most of the time, an annihilation event produces two gamma rays at 511 keV traveling at substantially 180 degrees apart.
By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can retrieve the likely location of the original disintegration. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated. In addition to the location of the two scintillation events, if accurate timing (within few hundred picoseconds) is available, a time-of-flight (TOF) calculation can add more information regarding the likely position of the event along the line. Limitations in the timing resolution of the scanner will determine the accuracy of the positioning along this line. Limitations in the determination of the location of the original scintillation events will determine the ultimate spatial resolution of the scanner, while the specific characteristics of the isotope (e.g., energy of the positron) will also contribute (via positron range and co-linearity of the two gamma rays) to the determination of the spatial resolution the specific agent.
The collection of a large number of events creates the necessary information for an image of an object to be estimated through tomographic reconstruction. Two detected events occurring at substantially the same time at corresponding detector elements form a line-of-response that can be histogrammed according to their geometric attributes to define projections, or sinograms to be reconstructed. Events can also be added to the image individually.
The fundamental element of the data collection and image reconstruction is therefore the LOR, which is the line traversing the system-patient aperture. Additional information can be obtained regarding the location of the event. First, it is known that, through sampling and reconstruction, the ability of the system to reconstruct or position a point is not space-invariant across the field of view, but is better in the center, slowly degrading toward the periphery. A point-spread-function (PSF) is typically used to characterize this behavior. Tools have been developed to incorporate the PSF into the reconstruction process. Second, the time-of-flight, or time differential between the arrival of the gamma ray on each detector involved in the detection of the pair, can be used to determine where along the LOR the event is more likely to have occurred.
The above described detection process must be repeated for a large number of annihilation events. While each imaging case must be analyzed to determine how many counts (i.e., paired events) are required to support the imaging task, current practice dictates that a typical 100-cm long, FDG (fluoro-deoxyglucose) study will accumulate several hundred million counts. The time required to accumulate this number of counts is determined by the injected dose of the agent and the sensitivity and counting capacity of the scanner.
PET imaging systems use detectors positioned across from one another to detect the gamma rays emitting from the object. Typically a ring of detectors is used in order to detect gamma rays coming from each angle. Thus, a PET scanner is typically substantially cylindrical to be able to capture as much radiation as possible, which should be, by definition, isotropic. The use of partial rings and rotation of the detector to capture missing angles is also possible, but these approaches have severe consequences for the overall sensitivity of the scanner. In a cylindrical geometry, in which all gamma rays included in a plane have a chance to interact with the detector, an increase in the axial dimension has a very beneficial effect on the sensitivity or ability to capture the radiation. Thus, the best design is that of a sphere, in which all gamma rays have the opportunity to be detected. Of course, for application to humans, the spherical design would have to be very large and thus very expensive. Accordingly, a cylindrical geometry, with the axial extent of the detector being a variable, is realistically the starting point of the design of a modern PET scanner.
Once the overall geometry of the PET scanner is known, another challenge is to arrange as much scintillating material as possible in the gamma ray paths to stop and convert as many gamma rays as possible into light. It is necessary to consider two dimensions of optimization in this process. On one hand, the “in-plane” sensitivity forces as much crystal as possible (crystal thickness) around the circumference of the detector. On the other hand, for a given crystal thickness, the axial length of the detector-cylinder will define the overall system sensitivity, which is roughly proportional to the square of the axial length (the solid angle subtended by a point in the middle of a cylinder). Practical cost considerations will unavoidably be part of the optimization process. Optimal distribution of the crystal and associated sensors is central to the overall system cost, as it typically represents up to two-thirds of the entire cost of the PET imaging system.
Conventionally, a cylindrical geometry is the design of choice for a PET scanner. As shown in FIG. 1, the cylindrical geometry can capture all events in the transaxial plane. The axial extent of the detector will determine how many such planes can be defined, as well as how many oblique planes can be utilized.
As shown in FIG. 1, the scanner is formed by a series of small blocks representing the detector elements. Only a few dozen detector elements are shown for simplicity. In reality, several hundred pixels are necessary to adequately sample the geometry. The same is true for the axial direction. Detector elements are typically the same size in both directions, but can also be of different sizes in the two dimensions. The cross-section of the patient shows a thorax, lungs, heart, and spine, wherein the patient is resting on the scanner patient pallet. The drawings further illustrates a few possible lines of response, representing positron annihilation events originating from the “heart” and being collected at various points on the scanner, both on the circumference and axially.
The overall dimension of the scanner typically varies from 70 to 90 cm in diameter so as to cover the whole human body. The axial dimension can vary more. Conventional scanners have at least 15 cm of axial coverage (to at least cover the heart), while larger dimensions are possible and desirable.
Further, current clinical practice places the patient more or less in the center of the scanner. Given that the patient typically comprises 50% or less of the scanner diameter, placing the patient slightly lower than the center is also desirable, providing more “breathing” space for the patient inside the scanner aperture.
Nevertheless, the goal of the PET scanner is to collect as many lines as possible from the patient in both the axial and transaxial planes.
While the conventional design provides an efficient geometry to collect positron annihilation events, it also defines a rigid set of rules for building the scanner, and consequently, offers few options for controlling its cost.
Attempts to increase the axial extent of the scanner have been proposed, but the rules for complete sampling in the transaxial plane have not changed.
Given the very high relative cost of the detector elements, any attempts to increase the scanner aperture (i.e., a larger diameter) or its axial extent require a significant cost increase. An increase in the diameter of the scanner is desirable to accommodate positioning (immobilization) tools for therapy (matching the radiation therapy unit), which require 85 cm or more, and also to provide better patient comfort by diminishing the claustrophobic stress still experienced by many patients. An increase in the axial extent of the scanner is desirable to increase the sensitivity (number of events being collected) and to cover larger organ or body sections. For example, the entire lung typically spans up to 25-30 cm and the head and neck requires at least 30 cm.
Thus, a real question facing scanner designers is, given a certain amount of detector material, which represents the predominant cost of the scanner, what geometry optimizes the number of event counts while providing adequate sampling for reconstruction.
Additionally, conventional designs do not provide a way to change or otherwise optimize the scanner for scanning a specific organ or region of interest in the patient. Rather, all images are acquired in the same way, regardless of the object or region of the object to be optimized.